The ability to deliver agents into and through skin surfaces (transdermal delivery) provides many advantages over oral or other parenteral delivery techniques. In particular, transdermal delivery provides a safe, convenient and noninvasive alternative to traditional drug administration systems, conveniently avoiding the major problems associated with oral delivery, e.g., variable rates of absorption and metabolism, gastrointestinal irritation and/or bitter or unpleasant drug tastes. Transdermal delivery also avoids problems associated with traditional needle and syringe delivery, e.g., needle pain, the risk of introducing infection to treated individuals, the risk of contamination or infection of health care workers caused by accidental needle-sticks and the disposal of used needles. In addition, such delivery affords a high degree of control over blood concentrations of administered drugs.
However, despite its clear advantages, transdermal drug delivery presents a number of its own inherent logistical problems. The passive delivery of drugs through intact skin necessarily entails the transport of molecules through a number of structurally different tissues, including the stratum corneum, the viable epidermis, the papillary dermis, and the capillary walls in order for the drug to gain entry into the blood or lymph system. Transdermal delivery systems must therefore be able to overcome the various resistances presented by each type of tissue. In light of the above, a number of alternatives to passive transdermal delivery have been developed. These alternatives include the use of skin penetration enhancing agents, or “permeation enhancers,” to increase skin permeability, as well as non-chemical modes such as the use of iontophoresis, electroporation or ultrasound. However, such techniques often give rise to unwanted side effects, such as skin irritation or sensitization. Thus, the number of drugs that can be safely and effectively administered using traditional transdermal delivery methods has remained limited.
More recently, a novel transdermal delivery system that entails the use of a needleless syringe to fire solid drug-containing particles in controlled doses into and through intact skin has been described. In particular, commonly owned U.S. Pat. No. 5,630,796 to Bellhouse et al. describes a needleless syringe that delivers pharmaceutical particles entrained in a supersonic gas flow. The needleless syringe is used for transdermal delivery of powdered drug compounds and compositions, for delivery of genetic material into living cells (e.g., gene therapy) or nucleic acid immunization, and for the delivery of biopharmaceuticals to skin, muscle, blood or lymph. The needleless syringe can also be used in conjunction with surgery to deliver drugs and biologics to organ surfaces, solid tumors and/or to surgical cavities (e.g., tumor beds or cavities after tumor resection). In theory, practically any pharmaceutical agent that can be prepared in a substantially solid, particulate form can be safely and easily delivered using such devices.
One particular needleless syringe generally comprises an elongate tubular nozzle having a rupturable membrane initially closing the passage through the nozzle and arranged substantially adjacent to the upstream end of the nozzle. Particles of a therapeutic agent to be delivered are disposed adjacent to the rupturable membrane and are delivered using an energizing means which applies a gaseous pressure to the upstream side of the membrane sufficient to burst the membrane and produce a supersonic gas flow (entraining the pharmaceutical particles) through the nozzle for delivery from the downstream end thereof. The particles can thus be delivered from the needleless syringe at delivery velocities as high as Mach 1 to Mach 8, which velocities are readily obtainable upon the bursting of the rupturable membrane.
Another needleless syringe configuration generally includes the same elements as described above, except that instead of having the pharmaceutical particles entrained within a gas flow, the downstream end of the nozzle is provided with a diaphragm which is moveable between a resting “inverted” position (in which the diaphragm presents a concavity on the downstream face to contain the pharmaceutical particles) and an “everted” position (in which the diaphragm is outwardly convex on the downstream face as a result of a supersonic shockwave having been applied to the upstream face of the diaphragm). In this manner, the pharmaceutical particles contained within the concavity of the diaphragm are expelled at a high initial velocity from the device for transdermal delivery thereof to a targeted tissue surface.
Transdermal delivery using the above-described needleless syringe configurations is carried out with particles having an approximate size that generally ranges between 0.1 and 250 μm. For drug delivery, a typical particle size is usually at least about 10 to 15 μm (the size of a typical cell). For gene delivery, a typical particle size is generally substantially smaller than 10 μm. Particles larger than about 250 μm can also be delivered from the device, with the upper limitation being the point at which the size of the particles would cause untoward damage to the skin cells. The actual distance which the delivered particles will penetrate depends upon particle size (e.g., the nominal particle diameter assuming a roughly spherical particle geometry), particle density, the initial velocity at which the particle impacts the skin surface, and the density and kinematic viscosity of the skin. In this regard, optimal particle densities for use in needleless injection generally range between about 0.1 and 25 g/cm3, preferably between about 0.8 and 1.5 g/cm3, and injection velocities generally range between about 150 and 3,000 m/sec.
A particularly unique feature of the needleless syringe is the ability to optimize the depth of penetration of delivered particles, thereby allowing for targeted administration of pharmaceuticals to various sites. For example, particle characteristics and/or device operating parameters can be selected to provide for penetration depths for, inter alia, epidermal or dermal delivery. One approach entails the selection of particle size, particle density and initial velocity to provide a momentum density (e.g., particle momentum divided by particle frontal area) of between about 2 and 10 kg/sec/m, and more preferably between about 4 and 7 kg/sec/m. Such control over momentum density allows for tissue-selective delivery of the pharmaceutical particles.
Accordingly, there is a need to provide a reliable method for preparing sufficiently dense particles (having a density of about 0.8 to 1.5 g/cm3) which have an average size of about 0.1 to 150 μm from a wide variety of pharmaceutical compositions. These pharmaceutical particles can thus be transdermally delivered to a subject using a needleless syringe system.
Needleless syringes, such as those described above, also provide a unique means for gene therapy and nucleic acid immunization. These techniques provide for the transfer of a desired gene into a subject with the subsequent in vivo expression thereof. Gene transfer can be accomplished by transfecting the subject's cells or tissues ex vivo and reintroducing the transformed material into the host. Alternatively, genes can be administered directly to the recipient.
A number of methods have been developed for gene delivery in these contexts. For example, viral-based systems using, e.g., retrovirus, adenovirus, and adeno-associated viral vectors, have been developed for gene delivery. However, these systems pose the risk of delivery of replication-competent viruses. Hence, nonviral methods for direct transfer of genes into recipient cells and tissues are desirable.
Nonviral methods of gene transfer often rely on mechanisms employed by mammalian cells for the uptake and intracellular transport of macromolecules. For example, receptor-mediated methods of gene transfer have been developed. The technique utilizes complexes between plasmid DNA and polypeptide ligands that can be recognized by cell surface receptors. However, data suggests that this method may permit only transient expression of genes and thus has only limited application.
Additionally, microinjection techniques have been developed for the direct injection of genetic material into cells. The technique, however, is laborious and requires single cell manipulations. Thus, the method is inappropriate for use on a large scale.
Direct injection of DNA-containing solutions into the interstitial space for subsequent uptake by cells has also been described. For example, International Publication No. WO 90/11092, published 4 Oct. 1990, describes the delivery of isolated polynucleotides to the interior of cells wherein the isolated polynucleotides are delivered into the interstitial space of the tissue and then taken up by individual cells to provide a therapeutic effect. Such methods entail the injection of the DNA-containing solutions into tissue using conventional needles or cannulas, and are therefore not well suited for long term therapies or for field or home applications.
Biolistic particle delivery systems (particle bombardment systems) have also been developed for gene delivery into plant cells. Such techniques use a “gene gun” to introduce DNA-coated microparticles, such as DNA-coated metals, into cells at high velocities. The coated metals (biolistic core carriers) are generally propelled into cells using an explosive burst of an inert gas such as helium. See, e.g., U.S. Pat. No. 5,100,792 to Sanford et al. The technique allows for the direct, intracellular delivery of small amounts of DNA.
Biolistic core carriers upon which the DNA is coated, such as tungsten, gold, platinum, ferrite, polystyrene or latex, have to date been needed to achieve adequate gene transfer frequency by such direct injection techniques. See, e.g., International Publication No. WO 94/23738, published Oct. 27, 1994. In particular, these materials have been selected based on their availability in defined particle sizes around 1 μm in diameter, as well as providing a sufficiently high density to achieve the momentum required for cell wall or cell membrane penetration. Additionally, common biolistic core carriers are chemically inert to reduce the likelihood of explosive oxidation of fine microprojectile powders, are non-reactive with DNA and other components of the precipitating mixes, and display low toxicity to target cells. See e.g., Particle Bombardment Technology for Gene Transfer, (1994) Yang, N. ed., Oxford University Press, New York, N.Y. pages 10-11.
However, such biolistic techniques are not appropriate for use with large DNA molecules since precipitation of such molecules onto core carriers can lead to unstable configurations which will not withstand the shear forces of gene gun delivery.
Accordingly, there remains a need to provide a highly efficient method for introducing therapeutically relevant DNA or other nucleic acid molecules into mammalian tissue cells wherein the method avoids the problems commonly encountered with prior gene delivery techniques.